Passive cardiac assistance device

ABSTRACT

Artificial implantable active and passive girdles include a heart assist system with an artificial myocardium employing a number of flexible, non-distensible tubes with the walls along their long axes connected in series to form a cuff and a passive girdle is wrapped around a heart muscle which has dilatation of a ventricle to conform to the size and shape of the heart and to constrain the dilatation during diastole. The passive girdle is formed of a material and structure that does not expand away from the heart but may, over an extended period of time be decreased in size as dilatation decreases.

[0001] This application is a continuation of U.S. patent applicationSer. No. 09/023,592 filed Feb. 13, 1998, which is a divisional of U.S.patent application Ser. No. 08/581,051 filed Dec. 29, 1995 (now U.S.Pat. No. 5,800,528 issued Sep. 1, 1998), which is a continuation-in-partof U.S. patent application Ser. No. 08/490,080 filed Jun. 13, 1995 (nowU.S. Pat. No. 5,713,954 issued Feb. 3, 1998). The contents of each ofthese patent applications are specifically incorporated herein byreference.

FIELD AND BACKGROUND OF THE INVENTION

[0002] The present invention relates to cardiac devices in general, andmore specifically to passive and active cardiac girdles.

[0003] Patients having a heart condition known as ventricular dilatationare in a clinically dangerous condition when the patients are in an endstage cardiac failure pattern. The ventricular dilatation increases theload on the heart (that is, it increases the oxygen consumption by theheart), while at the same time decreasing cardiac efficiency. Asignificant fraction of patients in congestive heart failure, includingthose who are not in immediate danger of death, lead very limited lives.This dilatation condition does not respond to current pharmacologicaltreatment. A small amount, typically less than 10%, of the energy andoxygen consumed by the heart, is used to do mechanical work. Thus thebalance, which is the major part of the energy consumed by the heart isused in maintaining the elastic tension of the heart muscles for aperiod of time. With a given pressure, the elastic tension is directlyproportional to the radius of curvature of the heart ventricle. Duringventricular dilatation the ventricular radius increases and the energydissipated by the heart muscle just to maintain this elastic tensionduring diastole is abnormally increased, thereby increasing oxygenconsumption.

[0004] A number of methods and devices have been employed to aid thepumping action of failing hearts. Many of these include sacs or wrapsplaced around the ailing heart, or, in some instances only around theventricle of the failing heart, with these wraps constructed to providefor active pumping usually, but not always, in synchronism with theventricular pumping of the natural heart. A number of cardiac assistsystems employing a variety of pumping approaches for assisting thepumping action of a failing natural heart have been developed. Thesesystems include those suitable for partial to full support of thenatural heart, short term (a few days) to long term (years), continuouspumping to various degrees of pulsability, and blood contacting versusnon-blood contacting. Table 1 lists a number of presently developeddevices with pertinent operating characteristics. TABLE 1 Level of BloodDevice Support Pulsatility Duration Contacting Comments IAPB Partial<20% Y Days to Months Y Counterpulsation provides LV unloading BiopumpFull N Days Y Limited to short duration due to thrombotic potentialThoratec Full Y Months Y Sac-type actuation Novacor Full Y Months YSac-type pump with electric actuation Hemopump Partial N Days Y Axialflow pump 50-75% Heart Mate Full Y Months Y Pusher-Plate pneumatic andelectric Aortic Patch Partial Y Months Y Counterpulsation BVS 5000 FullY Weeks Y Designed for temporary support Anstadt Full Y Days N Cardiacresuscitation Cardiomyoplasty Partial <20% Y Years N Requires muscletraining for active support

[0005] One, more recent development in the field of cardiomyoplastyinvolves the wrapping and pacing of a skeletal muscle around the heartto aid in the pumping. In that configuration, a pacemaker is implanted-to control the timing of the activation of the wrapped around skeletalmuscle.

[0006] A major consideration in the design of cardiac support systems isthe risk of thromboembolism. This risk is most associated with use ofartificial blood contacting surfaces. A variety of approaches have beenemployed to reduce or eliminate this problem. One approach has been theemployment of smooth surfaces to eliminate potential sites for thrombiand emboli generation as well as textured surfaces to promote cellgrowth and stabilization of biologic surfaces. One problem affectingthromboembolism risk in heart assists arises from the use of prosthetic,biologic or mechanical pericardial valves. This risk can be some whatlowered by the use of anticoagulation therapy. However, the use requirescareful manipulation of the coagulation system to maintain an acceptablebalance between bleeding and thromboembolic complications. The texturedsurface approach employs textured polyurethane surfaces and porcinevalves to promote pseudo-intima formation with a stable cellular lining.While thromboembolic rates resulting from these measures are acceptableas temporary measures, improvements, particularly for implantabledevices are highly desirable.

[0007] A second problem associated with implanted cardiac assist devicesis the problem of infection, particularly where the implanted device haslarge areas of material in contact with blood and tissue. More recentlyclinical protocols have improved and even the drive line and vent tubesassociated with implants that require some percutaneous attachments havebeen manageable. However, for a ventricular assist device, quality oflife considerations require that vent lines and drive lines which crossthe skin barrier be eliminated thereby avoiding the encumbrance topatient activities.

[0008] A third problem area in ventricular assist devices is thecalcification of these devices. This is particularly so for long termimplant situations which may last five years or more. Here again thecriticality of this factor is reduced for devices which do not involvedirect blood contact.

[0009] Another approach employed in ventricular assist has been thedevelopment of non-pulsatile pumps. However, once again, the blood isexposed to the surfaces of the pump, particularly the bearing and sealarea.

[0010] Unlike an entirely artificial heart, in which failure of thesystem leads to death, a ventricular assist device augments the impairedheart and stoppages should not result in death, unless the heart is incomplete failure. However, for most present ventricular assist devicesystems, stoppage of even a few minutes results in formation of bloodclots in the device, rendering any restart of the system a very riskyundertaking.

SUMMARY AND OBJECTS OF THE INVENTION

[0011] According to one aspect, the present invention an artificialmyocardium is constructed of an extremely pliant, non-distensible andthin material which can be wrapped around the ventricles of a natural,but diseased heart. This artificial myocardium mimics thecontraction-relaxation characteristics of the natural myocardium andprovides sufficient contractility, when actuated, to at least equal thecontractility of a healthy natural myocardium. In this arrangement allof the direct blood contact is with the interior surfaces of the naturalheart and surrounding blood vessel system. The device is hydraulicallyactuated in timed relationship to the contractions of the natural heart.

[0012] Using this system, the natural heart is left in place and theassist system supplies the reinforcing contractile forces required forsatisfactory ventricular ejection.

[0013] A key concept for this artificial myocardium system is achievedby the realization of a controllable, artificial myocardium employing acuff formed of a series of closed tubes connected along their axiallyextending walls. With sufficient hardware to hydraulically (orpneumatically) inflate and deflate these tubes, a controlled contractionis produced as a result of the geometric relationship between the lengthof these series of tubes in deflated condition and the length of theseries of tubes when they are fluidically filled in the inflatedcondition. If the cuff is formed of a series of “n” tubes, each ofdiameter “d” when inflated, connected in series, the total perimeterlength of this cuff when deflated is given by n(πd/2). However, whenthese tubes are filled with fluid, they have a circular cross-sectionsuch that the length of the cuff is the sum of the diameters in theindividual tubes or nd. Thus the ratio of the change in perimeter lengthbetween the collapsed and the filled state is π/2. If this cuff iswrapped around the natural heart, it will, when pressurized, shorten andsqueeze the heart by producing a “diastolic” to “systolic” length changeof 36%. Typical sarcomere length changes are approximately 20%.

[0014] Suitable hardware, including a hydraulic pump, a compliantreservoir and rotary mechanical valve, together with appropriateactuating electronics can all be implanted in the patient's body. If thepower source is an internal battery, then power may be transcutaneouslytransmitted into the body to recharge this battery.

[0015] Ventricular dilatation is a clinically dangerous condition forend stage cardiac failure patients. The output of the heart is effectedby: (a) end-diastolic volume (ventricular volume at the end of thefilling phase), (b) end-systolic volume (ventricular volume at the endof the ejection phase), and (c) heart rate. When (a) is very large, (b)also tends to be larger and (c) tends to be larger than normal. Allthree of these factors contribute to large increases in the tension-timeintegral and therefore to increased oxygen consumption.

[0016] Only a small amount of the energy consumed by the heart is usedto do mechanical work. For example, with a cardiac output of 5liters/minute, and Δp of 100 mm(Hg), the mechanical work done by theleft ventricle is about 1.1 watts, and that of the right ventricle isabout 0.2 watts. This compares with the typical total energy consumed bythe heart (mechanical work during systole plus the energy cost inmaintaining elastic tension during diastole) of about 12 to 15 watts.

[0017] Thus, since cardiac efficiency (typically between 3% and 15%) isdefined as the ratio of the mechanical work done by the heart to thetotal energy (or load of the heart muscle): then,

[0018] Cardiac Efficiency,$\eta = \frac{\int{P_{v}{V}}}{{\int{P{V}}} + {k{\int{T{r}}}}}$

[0019] where

[0020] P_(v): Ventricle Pressure

[0021] P: Pressure

[0022] V: Volume

[0023] T: Tension

[0024] t: Time

[0025] The constant k accounts for conversion of units.

[0026] An increase in mechanical work by a large factor results in asmall increase in oxygen consumption but an increase in tension timecauses a large increase in oxygen consumption. Patients with dilatedventricles who have undergone active cardiomyoplasty have not beenreported to show any objectively measurable hemodynamic improvement.

[0027] According to a further aspect of the present invention, acompletely passive girdle is wrapped around the ventricle or the entireheart muscle, and sized so that it constrains the dilatation duringdiastole and does not effect the action of the ventricle during systole.With the present surgical techniques, it is expected initial access tothe heart to place the girdle in position, will require opening thechest. However, it may be possible to locate a girdle in positionwithout thoracotomy. In one embodiment, a synthetic girdle made frommaterial that can limit tension, but is otherwise deformable to conformto the anatomical geometry of the recipient heart is used. This girdlemay be adjustable in size and shape over an extended period of time inorder to gradually decrease the ventricular dilatation. A secondembodiment employs a fluid filled passive wrap constructed of a seriesof horizontal sections. This provides for a variable volume to beenclosed by the wrap with volume control being obtained by controllingthe volume of fluid from an implantable reservoir within the body. Inits most preferable form, this passive wrap can be formed of a series ofhorizontal tubular segments each individually sealed and attached to oneanother along the long axis of the cylinder. If the cylinders are madeof indistensible material, then changing the volume of fluid from thecylinders being in a substantially deflated condition to one where theyare partially or fully inflated, decreases the internal perimeter of thewrap or girdle, thereby decreasing the effective radius of the girdlearound the heart. Another feature of the invention is a feedback system,wherein sensors, for example, strain gauges, can be built into anindistensible lining to measure its tension and thereby provideautomatic feedback to a hydraulic circuit controlling the wrap volume.

[0028] To avoid the problem of potential irritability and damage to theexternal myocardium cells by virtue of the artificial wrap and its longterm constraining contact with the myocardium, one embodiment of theinvention employs a tissue engineered lining to protect the myocardium.This tissue engineered lining consists of a polymer scaffold seeded withmyocardial cells harvested from the patient's own myocardium usingtissue engineering technology. That lining then generates a biologicalmyocardio-interfacing surface and remains firmly attached to the polymerinterfacing with the surface from which the wrap is made. Such a liningwould integrate biologically to the heart's myocardial cells in a manneranalogous to other devices currently being investigated which use cellscaffolds for in vitro and in vitro tissue engineering.

[0029] It is therefore an object of the present invention to produce aventricular assist device system employing an artificial myocardiumplaced around the natural heart (extra cardiac assist). This design,then, does not contact the bloodstream eliminating many of the problemsdiscussed above.

[0030] It is another object of this invention to provide a ventricularassist device which mimics the action of the natural heart whileavoiding the compressive action of the direct mechanical ventricularactuation systems on the epicardium.

[0031] It is a further object of this invention to provide an artificialmyocardium in which the external fluid being pumped is a fraction of theblood volume pumped by the action of the artificial myocardium.

[0032] It is yet another object of this invention to provide aventricular assist device which is compact, requires relatively lowenergy input and does not require percutaneous components.

[0033] It is a further object of this invention to provide a passivegirdle to be wrapped around a heart suffering from ventriculardilatation to limit this dilatation and thus improve the performancecharacteristics of the heart.

[0034] It is another object of this invention to provide a passivegirdle or vest which can, over a period of time, have its diameterdecreased to effect some decrease in dilatation of the ventricle.

[0035] Other objects will become apparent in accordance with thedescription of the preferred embodiments below.

BRIEF DESCRIPTION OF THE DRAWING

[0036]FIGS. 1A and 1B are diagrammatic illustrations of the tubeconstruction of an artificial myocardium in accordance with theprinciples of this invention;

[0037]FIG. 2 is a generally diagrammatic illustration of the anatomicalplacement of the cardiac assist system of this invention in the humanbody;

[0038]FIG. 3A is an illustration in diagrammatic form of the perimeterlength of the artificial cuff at various stages of the natural heartcontractions;

[0039]FIG. 3B is an illustration of a cross-sectional view of thesystolic and diastolic shapes of the artificial myocardium;

[0040]FIG. 4A is a diagrammatic view of a partial wraparound of the leftventricle by the artificial myocardium in the systolic state and FIG. 4Bis a diagrammatic view of a partial wraparound of the left ventricle bythe artificial myocardium in the diastolic state;

[0041]FIG. 5 is a diagrammatic illustration of the geometricrelationship of the partially inflated tubes and the encircled heartrepresented by a radius of D/2.

[0042]FIGS. 6A, 6B and 6C illustrate factors involved in altering thelength of the individual tubes when inflated or deflated, to achievespecific shrink ratios.

[0043]FIG. 7 is a graphical representation of the ratio of hydraulic toafterload pressure as a function of the tube inflation parameter, θ, forn=10;

[0044]FIG. 8 is a graphical representation of the timing relationshipfor device actuation;

[0045]FIG. 9 is a graphical representation of the left ventricularpressure versus the left ventricular volume under various conditions;

[0046]FIG. 10 is a block diagram of the hydraulic system of theartificial myocardium of this invention;

[0047]FIG. 11A is a cross-sectional view of an energy converter andvalve structure in one position for use in the artificial myocardium ofthis invention;

[0048]FIG. 11B is a cross sectional view of the same structure as FIG.11A but with the valving in a different position for pumping from thehydraulic cuff to the hydraulic reservoir;

[0049]FIG. 12A is a cross-sectional view across the skin interface of asubdermal port for emergency access and manual pumping;

[0050]FIG. 12B is an “X-ray” view of the access system of FIG. 12Aviewed externally;

[0051]FIG. 13 is a graphical representation of the hydraulic tophysiological pressure ratios as a function of the number of tubes forboth univentricular and biventricular support;

[0052]FIG. 14 is a graphical representation of the physiological tohydraulic volume ratios as a function of the number of tubes for bothuniventricular and biventricular support;

[0053]FIG. 15 is a graphical representation of the blood stroke volumeas a function of the number of tube segments in the artificialmyocardium of this invention;

[0054]FIG. 16 is a diagrammatic illustration of a mock loop used for invitro tests;

[0055]FIG. 17 is a graphical representation of the relationship betweenAfterload Pressure (AOP) and Driving Pressures (P-Drive) for theartificial myocardium; and

[0056]FIG. 18 is a graphical representation of flow sensitivity to drivepressure at a constant afterload pressure;

[0057]FIG. 19A is an illustration generally in cross sectional form of aheart girdle constructed in accordance with the principles of thisinvention;

[0058]FIG. 19B is an illustration in cross-sectional form of the heartgirdle of FIG. 19A with the girdle in a pneumatically filled condition;

[0059]FIG. 20 is a perspective view of the heart girdle of FIGS. 19A and19B showing the horizontal segments.

[0060]FIG. 21 is an illustration generally in block diagram form of acontrol system for the heart girdle of FIG. 19 including a strain gaugeand electronic actuator to maintain constant tension at the interfacebetween the girdle and the heart muscle;

[0061]FIG. 22 is an illustration in perspective view of a heart girdleemploying a flexible mesh of interlocked circular plastic loops;

[0062]FIG. 23 is an illustration of a portion of a girdle constructedgenerally in accordance with the girdle construction of FIG. 22, butfurther including strings adapted to draw the girdle into decreasingdiameter shape;

[0063]FIG. 24 is another embodiment of a portion of a passive girdleformed of a material characterized by a specific internal structure; and

[0064]FIG. 25 is a cross sectional drawing of a girdle-myocardiuminterface constructed of biologically engineered myocardial tissue.

DESCRIPTION OF PREFERRED EMBODIMENTS

[0065]FIG. 1A and FIG. 1B illustrate diagrammatically the operation ofthe artificial myocardium. The artificial myocardium 11 is formed of aseries of tubes planed together in series to form, in this instance acomplete circle, which in FIG. 1A has a diameter D_(d). In FIG. 1B thefeature of the tubes is filled hydraulic fluid producing a circularcross-section, shortening the total perimeter of the circular cuff to acircle having a diameter D_(S). Referring to FIG. 1B, if the diameter ofthe tube with the circular cross-section is d, then the diameter of thecircular cuff is approximately equal to nd/π, where n equals the totalnumber of tubes. On the other hand, when the tubes are no longer filledwith hydraulic fluid and are collapsed then the diameter D_(d) isapproximately equal to$\frac{n\left( {\pi \quad {d/2}} \right)}{\pi}$

[0066] These expressions follow from the consideration that the seriesof n tubes in the inflated condition, as illustrated in FIG. 1B form acircle with the number of tubes times diameter of each of the individualtubes. On the other hand in the collapsed condition each one of thetubes has a length equal to its perimeter divided by 2. Since theperimeter is πd then the length of each collapsed tube is πd /2 and thediameter of the cuff in this condition is the sum of the length of thecollapsed tubes over π.

[0067] When this cuff is placed around a natural heart and the fillingand emptying of the tubes is in phase with the systole and diastole ofthe natural heart, then the shortening of the cuff forces the surroundedventricle to decrease its diameter thereby causing the ventricle toeject blood. The ejection fraction of this artificial myocardium isindependent of the number of tubes or the heart dimensions. The ejectionfraction is a function of only the hydraulic pressure. When thehydraulic pressure is large enough to inflate the tubes to cylinders,the ejection fraction is,$E_{f} = \frac{D_{d}^{2} - D_{S}^{2}}{\left( {D_{d} - {2t}} \right)^{2}}$where${D_{S} \cong \frac{nd}{\pi}},{D_{d} = \frac{nd}{2}},{t = {{heart}\quad {muscle}\quad {thickness}}}$

[0068]FIG. 2 illustrates an artificial myocardial assist system locatedin the human body. In the illustrated system the artificial myocardium11 is shown as a cuff placed around the ventricles of the natural heart17. The hydraulic fluid is pressurized by energy converter 19 either inthe direction of the cuff 11 or of a compliant hydraulic reservoir 21.The energy converter 19 is electrically controlled by virtue of internalelectric circuit 23 which is powered by an internal battery 25. Theinternal electrical circuit 23 is also coupled to external battery 27via a transcutaneous electrical terminal (TET) 31. The energy converter19 consists of a hydraulic pump coupled to a brushless electric motor toshuttle fluid between the artificial myocardium 11 and the compliantreservoir 21. Flow switching is accomplished by a rotary mechanicalvalve incorporated into the energy converter, which in turn issynchronized by a control signal generated by detection of the R wavefrom the ECG signal in the natural heart. Continuous adjustment of thehydraulic pump output allows the level of cardiac assist to be varied ona beat-by-beat basis.

[0069]FIG. 3A is a graphical illustration of the length of the perimeterof the artificial myocardium during systole, mid systole and diastole.FIG. 3B is an illustration in cross section view of the systolic anddiastolic shapes of the artificial myocardium in a cylindrical geometry.The outer ring illustrates the cylindrical tubes in collapsed form,while the inner ring illustrates those same tubes when they are filledwith hydraulic fluid during the systole. The natural heart pumps bloodprimarily through circumferential contraction. Most of the diastolic tosystolic volume change is derived primarily from the 20% change in thecircumference component and to a lesser extent the 9% change in theaxial length. As can be seen in FIG. 3A and FIG. 3B, the volumetricchange of the myocardium is 36% from the relaxed (diastole) position tothe fully contracted (systole) position. In the cross-sectional view,and assuming that the artificial myocardium were a completelycylindrical cuff, there is a 60% change in the area between these statesin the artificial myocardium, equivalent to a 60% ejection fraction of ahealthy heart. Although the description is based on a cylindricalgeometry, with interconnecting tapered tubes, the artificial myocardiumwill match the conical shape of the heart when appropriate taper angleis selected for the tubes.

[0070] With this hydraulic design, the natural heart having a typicalmyocardium thickness, a heart base diameter of 80 mm and an axialventricular length from apex to base of 50 mm, a left ventricular wrapof the artificial myocardium results in a stroke volume of 83 cc. Thesevalues are the same as that which would be expected from a normallyoperating left ventricle.

[0071] One very important factor in the operation of the artificialmyocardium is that the hydraulic pressure required for contractionagainst a given ventricular pressure is directly proportional to thenumber of tubes n forming this artificial myocardium. From energyconservation principles, the hydraulic flow in this artificialmyocardium varies inversely with the number of tubes. For example, withtypical natural heart dimensions, and a hydraulic stroke volume of 24cc, a pressure of 760 mm Hg produces a left ventricular stroke volume of95 cc at a mean aortic pressure of 90 mm Hg. Importantly, the hydraulicflows required are much less than the generated blood flow. In thisexample, the hydraulic flow is approximately 25% of the blood flowproduced. These smaller hydraulic flows result in lower hydraulic lossesand higher efficiencies. This, taken together with the smallerdimensions for the energy converter and the compliance chamber is veryadvantageous for an implanted device.

[0072] The output of a single energy converter can simultaneouslygenerate different contractile forces for left and right ventricleassist by varying the number of tubes which are wrapped around the leftand right ventricles while maintaining an equal drive pressure. If theartificial myocardium on the right side has M times the number of tubesas that on the left side, the contractile pressures on the right sidewill be M times lower. With this arrangement, the artificial myocardiummay be tailored to match differing afterloads from the two ventricles. Agood design parameter for considerations for efficiency and tubedynamics would be to operate at a contraction of 22%, which is veryclose to the contraction value of a natural healthy myocardium. It isbelieved that with this contraction level, the low mechanical stresseson the artificial myocardium may well result in an operation life offive years, a high reliability for the artificial myocardium.

[0073] In the artificial myocardium assist system of this embodiment thecontribution from the artificial myocardium is additive to the naturalheart with a timing cycle synchronized with the ECG, so that the controlalgorithm can adjust the hydraulic flow on a beat-by-beat basis toachieve the desired ejection fraction. Thus, if the natural myocardiumwas completely healthy, minimal pumping would be required by thehydraulic pump. On the other hand, if the natural heart had very littlemyocardial contractility, then the artificial myocardium would providealmost the entire contracting force.

[0074] With this design of the artificial ventricles there is no bloodcontact with the artificial surfaces of the cardiac assist system,thereby avoiding the principal concerns of the thromboembolism risk.Another important safety consideration, is that if the artificialmyocardium system were to stop, the natural forces on the hydraulicfluid will cause it to empty from the artificial myocardium and flowinto the compliant reservoir. The only effect of these conditions on thecardiovascular system would be those caused by the collapsed flexiblewrap on the myocardium. A subdermal port could be provided to allowemergency actuation of the artificial myocardium with a pneumatic pumpplaced external to the patient's body. Another consideration is that thecontrol algorithm of the artificial myocardium assist system can bearranged to provide contractions at a fixed predetermined rate if thereshould be a ventricular fibrillation or tachycardia of the naturalheart.

[0075] Because the design provides for no contact between the blood andthe components of the artificial myocardium assist system, thebiocomparability factors are limited to those which relate to theinterface between tissues and these components. In this embodiment, thetissue contacting material of the artificial myocardium may be apolyetherurethane, Angioflex®, manufactured by ABIOMED, Inc. of Danvers,Mass.

[0076]FIGS. 4A and 4B illustrate an example of a partial wrap around theheart for left ventricular support. The full wrap can be used forbiventricular support. FIGS. 4A and 4B show the systolic and diastolicpositions of the artificial myocardium under this condition. Each of then tubes are attached at their outer wall to two neighboring tubes exceptfor the two ends. Each tube, when inflated, has a diameter d, resultingin a wrap length of nd. Conversely, when the tubes are deflated the wraplength is

nπd/2

[0077] As illustrated in FIG. 5, when the tubes are partially inflatedthe length of the wrap is given by

L=ndπsinθ/,2θ

[0078] where θ is the angle representative of the curvature of the wallof the partially inflated tube and is defined as

θπd/4r

[0079] where r is the radius of the arc of each half of the tube wallwhen inflated.

[0080] Although the ventricle is conical in shape and accordingly theartificial myocardium is conformed to that shape, for simplicity therepresentation in FIGS. 4A and 4B is of a cylindrical shape. In FIG. 3Bthe effects of both tube inflation and stroke volume are shown. Thesystolic contracted shape 13 a of the artificial ventricle is plottedconcentrically inside the diastolic distended shape 13 b. The shadedannulus portion 15 of FIG. 3B represents the stroke volume change due tocontraction, while the dotted circle enclosed portion of the inflatedtubes represents displacement volume change. For the artificialmyocardium this displacement volume has a minor contribution to thestroke volume.

[0081] The volume may be expressed as$S_{V} = {{\frac{\pi}{4}{D_{d}^{2}\left\lbrack {1 - \left( {1 - \chi + {\chi \frac{\sin \quad \theta}{\theta}}} \right)^{2}} \right\rbrack}l} + {\frac{1}{2}V_{H}}}$

[0082] where l is the perimeter length of the cuff, where D_(d) is, asshown, the diastolic diameter of the ventricle and χ is the fraction ofthe ventricle being wrapped. V_(H) represents the actual displacementdue to the effects of tube inflation in addition to the stroke volumederived from the contraction. If values are substituted in thisequation, assuming a typical natural heart diameter of approximately 8cm, a length approximately 5 cm and a left side partial wrap of χ=½, thecontractile change from the uninflated to fully inflated tubes resultsin a stroke volume of approximately 83 cc. V_(H) is assumed to be zero.Similarly for a full wrap, the stroke volume would be approximately 150cc, equivalent to the sum of left and right side stroke volumes.

[0083] The ejection fraction achievable using such an assist device maybe estimated. The diastolic ventricular volume is given by${V_{d} = {\frac{\pi}{4}\left( {D_{d} - {2t_{d}}} \right)^{2}l}},$

[0084] where t_(d) is the ventricular wall thickness. Since the ejectionfraction is defined as the stroke volume divided by the diastolicventricular volume, it is given by${{EF} = \frac{SV}{\frac{\pi}{4}\left( {D_{d} - {2t_{d}}} \right)^{2}l}},$

[0085] Thus, the maximum ejection fraction obtainable for a typical wallthickness of 1 cm and a D_(d) equal to 8 cm is 59%, a number which isconsistent with the ejection fraction for a normal healthy heart.

[0086]FIG. 6A describes a section of a tube having an inflated diameterd and an inflated length l. The tube ends are formed to hemispheres inthe inflated position, the hemisphere radius being d/2.

[0087] The following relation can be written:

l _(i) =l _(o) +d  (1)

[0088] where:

[0089] l_(i) is the inflated length of the tube

[0090] l_(o) is the straight part of the tube

[0091] d is the inflated diameter of the tube.

[0092]FIG. 6B describes the tube when it is deflated. The flatteneddiameter of the tube is Bd/2 and its flattened length is:

l _(d) =l _(o) +Bd/2  (2)

[0093] where:

[0094] l_(d) is the deflated length of the tube.

[0095] The shrink ratio, defined as the change in linear dimensionsrelative to the original linear dimension l_(d), is: $\begin{matrix}{R = {\frac{\left\lbrack {l_{o} + \frac{\pi \quad d}{2}} \right\rbrack - \left\lbrack {l_{o} + d} \right\rbrack}{l_{o} + \frac{\pi \quad d}{2}} = {\frac{d\left\lbrack {\frac{\pi}{2} - 1} \right\rbrack}{l_{o} + \frac{\pi \quad d}{2}} = \frac{\pi - 2}{{2\frac{l_{o}}{d}} + \pi}}}} & (3)\end{matrix}$

[0096] Where:

[0097] R=relative change in linear dimensions.

[0098] As can be seen from the equation, the ratio R is only a functionof the ratio between the diameter d and the length l_(o). It can also beshown that this ratio does not change if a number of these tubes areconnected in series provided l_(o)/d remains unchanged.

[0099] Example: If the required longitudinal shrinkage is 12%, the ratiol_(o)/d can be calculated, using equation (3) to be:$0.12 = \frac{\pi - 2}{\frac{{2l_{o}} + \pi}{d}}$

[0100] To achieve a contraction of 12% longitudinally, the tube lengthl_(o) should be 3.18 times the diameter d. In a particular case, wherethe tube diameter is 10 mm, the overall tube length will be 41.8 mm.This length is about half of the required length and therefore, two ofthese tubes will be connected in series to achieve this goal. The seriesconnection can be done by making two wraps of half length, or, makingthe individual tubes with a shape as described in FIG. 6C.

[0101] As discussed earlier the contractile action of the artificialmyocardium results from the inflation of the series of tubes that arephysically attached to each other. The inflation of these tubes can beaccomplished either pneumatically or hydraulically. For a permanentlyimplantable device, the hydraulic approach is more practical. With apneumatic system, even in the absence of leakage losses, gas permeationacross flexing membranes is unavoidable. This effect is not probable ina hydraulic system with a proper choice of working fluids. In addition,the hydraulic system is safer in the event of rupture failure since highpressure cannot be maintained in the event of a leak. In the case of apneumatic system, a severe leak could result in cardiac compression. Inthe situation where the artificial myocardium is being employed as anassist device, the natural myocardium can generate some tension, and theartificial myocardium need only generate sufficient tension to boost theintraventricular pressure. Accordingly, the differential pressure, Prequired of the artificial myocardium may be only 20-30 mm Hg, boostingthe ventricular pressure from, for example, 60 to 80-90 mm Hg. The ratioof hydraulic pressure P_(H) to the load pressure P is expressed by$\frac{P_{H}}{P} = {\frac{n\quad \tan \quad \theta}{2\pi \quad \cos \quad \alpha} + {\frac{1}{2}\left( {1 - {\tan \quad \theta \quad \tan \quad \alpha}} \right)}}$

[0102] where α is the half angle subtended by each tube centered to theventricular axis.

[0103] Accordingly, the pump pressure P_(H) is related to threeparameters. First it is directly proportional to load pressure. Thushigher load pressures requires higher drive pressures. Secondly, as thenumber of tubes, n increases, the drive pressure required for a givenboost of intraventricular pressure increases. This is accompanied by aconcomitant decrease in the fluid flow of volume per stroke required ofthe hydraulic pump system. Third, the hydraulic pressure requiredincreases as the tubes are progressively inflated from being fullydeflated to fully circular.

[0104] Illustrated in FIG. 7 is the ratio of the hydraulic to theafterload pressure as a function of the parameter of tube inflationangle 2 for n=10. In the curve of FIG. 7, the targeted operating pointis shown by the asterisk. As can be seen a significant contraction isachieved when theta is between 1.2 to 1.4 radians, representing a 22% to30% contraction. Thus, the operational range of the hydraulic pressureis near one atmosphere for full assist against an afterload pressure ofapproximately 90 mm Hg, and ¼ atmosphere for cardiac boosting, that is,increasing the ventricular pressure by 20 to 30 mm Hg.

[0105] As discussed earlier the artificial myocardium requiressynchronization of its contraction with the natural heart. Contractionof the device must be timed appropriately with the heart's systole.Additionally, the drive pressures during systolic ejection must bemaintained to match the needs of physiologic afterloads.

[0106] The first factor, that is, the timing, can be achieved byimplanting an epicardial lead in a myocardial region of the naturalheart not in contact with the artificial myocardium, which could be nearthe apical region, or at the right atrial appendage. The artificialmyocardium would be timed to contract with the R wave produced on thislead. This detection can employ hardware that is used at the presenttime in implantable defibrillators. The systolic duration (inmilliseconds) is preprogrammed to match its functional dependence to thebeat rate (BR) in beats per minute expressed as τ_(s)=549 msec-2×BR.

[0107] In FIG. 8 the timing relationship between the ECG and theartificial myocardium is represented graphically. τ_(R) is the delaytime between the start of the artificial myocardium systole relative tothe R wave. τ_(S) is the artificial myocardium systolic duration, andτ_(d) is its diastolic duration. The exact coincidence of the start ofthe diastolic duration and the ECG T-wave is not critical. Largedeviations from this coincidence could either provide insufficientsupport (τ_(d) starts too early) or hamper diastolic filling (τ_(d)starts too late). Of course, irregular rhythms in the natural heart,such as the occurrence of bigeminy, premature ventricular contractions(PVC), or transient arrhythmias can also affect performance of theartificial myocardium system. The most straightforward way of dealingwith this situation is to cause the artificial myocardium wrap to beimmediately deflated to assume the diastolic state when this occurs. Itcan also be arranged so that when no ECG trace is detected, the devicewould contract at beat rates consistent with maintaining physiologicfilling pressures. Thus in the extra-cardiac support system of thisinvention certain unique operating characteristics can be provided whenworking in conjunction with the natural heart. It can, for example,provide extra contractility not present in other support approaches.

[0108]FIG. 9 illustrates the left ventricular pressure/volumerelationship. In FIG. 9 the left ventricular pressure is plotted againstthe left ventricular volume. The solid curve (loop 1) illustrates theperformance of the natural heart, while the dotted curves show how thepressure volume relationship may be altered in the presence of the extracardiac assist device. There are two factors that change as a result ofthe support device: the systolic pressure and the stroke volume. Theoptimal assist mechanism for this device is to boost the systolicpressure while allowing the myocardium to retain its isovolumetriccharacteristics by elevating the epicardial pressure (P_(e)) andincreasing transmyocardial pressure (P_(t)). The net effect is todisplace the isovolumetric curve upwards by P_(e). This is illustratedby the dashed line A shown in FIG. 9. Whether the intraventricularpressure (P_(V)) remains at P_(t), achieves the maximum value ofP_(t)+P_(e), or more likely reaches somewhere in between the twoextremes, depends on the vascular resistance and the ventricular strokevolume. FIG. 9 illustrates the three possibilities. In loop 2, theelevated ventricular pressure is matched by an increase in theafterload. This results in no change in stroke volume. In loop 3, theafterload remains unchanged, while the stroke volume is increased by theassist device. Case 2 would result if the ventricular stroke volume isequal to or greater than that available from the device. Thisrepresents, nominally, a healthy heart which requires no assist. Thecontrol scheme will be based on achieving a full systolic contractioneven though the diastolic ventricular filling may not be complete on abeat by beat basis.

[0109] With this control scheme, for a healthy heart which can generatesufficient tension against physiologic afterloads, the contractilityrequired of the artificial myocardium would be zero, and minimalhydraulic power would be required to fill the tubes. This in turngenerates minimal epicardial pressures and the assist is at a minimum.However, in cases where the natural heart is not capable of generatingnormal ejection fractions, the device contraction will extend the strokevolume of the ventricle. In order to realize these higher stroke volumesfrom the ventricle, the artificial myocardium must provide theadditional contractility needed. This requires higher hydraulic powerresulting in higher epicardial pressure translating to higherventricular pressure. Under such circumstances, assist will increaseflow and the afterload will also increase as a result of the increasedflow. This is illustrated by curve 4, the most likely operating P-V loopunder assist conditions.

[0110] The control scheme is relatively straightforward. The device willbe operating at a full systolic stroke in every beat. Power required toachieve the full stroke will be adjusted on a beat by beat basis. Thehydraulic stroke volume will be measured on every beat in order topermit implementation of this algorithm. During diastole, hydraulicpressure will be measured to provide an indication of the end diastolicpressures. This information will be used to determine system beat ratefor a heart with no rhythm.

[0111]FIG. 10 illustrates in block diagrammatic form the hydraulicsystem including the artificial myocardium 11, the energy converter 19and the hydraulic reservoir 21. In one embodiment, the artificialmyocardium 11 consists of four layers of polyetherurethane (Angioflex®)reinforced by polyester mesh, fabricated from 40 micrometer fibers. Twolayers of the fiber are interwoven to yield tubular interconnections.Separators are inserted in the individual tubes to prevent the opposinginner walls from bonding with each other during the manufacturingprocess. The process yields a device that has an overall wall thicknessof approximately ½ mm and a strength of 150 lbs/in, nearly two orders ofmagnitude higher than the tensile forces experienced by the device.Attached to the epicardium of the natural heart, the device, whendeflated, represents no additional diastolic resistance to the heart.The energy converter 19 shuttles fluid between the flexible reservoir 21and the artificial myocardium 11. Fluid reversal is achieved by arotating valve. The system flow resistance is designed to be low suchthat in the event of a stoppage of the system, fluid in the tubes of theartificial myocardium will automatically empty into the reservoir 21within a few heart beats. The positive diastolic filling pressure of thenatural hem and the negative intrathoracic pressure insures a drivingforce to empty the fluid from the device. Once it is completely emptiedthe artificial myocardium becomes a highly flexible sheet which followsthe wall motion of the natural hem without any additional resistance.Such a device can be restarted without any fear of embolic complicationsafter a temporary stoppage.

[0112] The direction of hydraulic fluid flow during systole in thiscardiac assist device is from the compliant fluid reservoir, through thedistributing manifold, and into the individual tubules of the artificialmyocardium. During diastole these bladders must be emptied by reversingthe direction of hydraulic fluid flow and pumping fluid back into thereservoir. The direction of fluid movement will be reversed by a rotaryporting valve in conjunction with unidirectional operation of thecentrifugal pump itself.

[0113] Unidirectional pump operation has several advantages overreversing the pump direction. Principally, unidirectional rotation ofthe pump shaft presents much more favorable conditions for bearing life.Although unidirectional pump speed will most likely change betweensystole (artificial muscle inflation) and diastole, these accelerationsand deceleration represents a fraction of those that would be associatedwith complete reversal of pump direction. Furthermore, in aunidirectional mode, pump impeller design can be optimized for fluidmotion in one direction.

[0114] Some of the key dimensions and motor performance parameters for asuitable motor for this system are listed below:

[0115] overall diameter: OD=1.0 in

[0116] overall length: OL=0.5 in

[0117] torque constant: KT=0.4 oz-in/amp

[0118] voltage constant: K_(B)=0.3 V/kRPM

[0119] terminal resistance: R_(M)=0.325 ohms

[0120] viscous damping: FI=0.002 oz-in/IRPM

[0121] The motor performance parameters listed above were used topredict the anticipated torque-speed and efficiency characteristics ofthe motor. The torque-speed performance characteristics indicate thatthe specific operating point of 1.2 oz-in at 35000 RPM can be obtainedwith an applied voltage of 12 V. In this application, the supply voltagewill be larger than 12 V, therefore pulse width modulation (PWM) of themotor supply can and will be used to control speed.

[0122] Unidirectional pump operation implies that fluid flow reversalwill be accomplished using a rotary porting valve. As illustrated inFIGS. 11A and 11B, a balanced inflow and outflow porting configurationis designed to minimize radial loads. The rotary porting valve 32consists of two concentric sleeves 34 a and 34 b, a fixed inner sleeve(not shown) and an outer sleeve 34 a which can be rotated through adefined angle by a torque motor (not shown). The valve is composed oftwo pairs of inlet ports leading to the impeller intake, one pair 40coming from the fluid reservoir and the other pair 38 from the hydrauliccuff 11, and two corresponding pairs of outlet ports off the impeller,one leading to the distributing manifold and one to the fluid reservoir.Accordingly, the valve is designed so that switching the outer sleeveinto the systolic position opens the inlet port from the reservoir whileclosing that from the manifold, and opens the outlet port to themanifold while closing that to the reservoir. Conversely, switching theouter sleeve back reverses the inlet and outlet ports and generatesdiastole. The rotary porting valve also incorporates hydraulic dampersto prevent valve rebound at the end of switching. FIGS. 11A and 11Billustrate this valving scheme.

[0123] The pump motor bearing is a component which requires carefuldesign considerations By the choice of a unidirectional pump, the majorfailure mode of bearings due to pump reversals inherent in some designshas been eliminated. However, the artificial myocardium operates at ahigh RPM in the range of 5,000 to 35,000 RPM. This high RPM places astringent requirement on the motor bearings.

[0124] Bearing load reduction can be achieved by judicious design.Although the outflow pressure is high in the AMA, the surface areas ofthe energy converter housing exposed to the pressure difference isrelatively small, such that the load on the bearings generally remainsin the same range. The use of a symmetric paired porting design for theinflow and outflow orifices, allows the radial loads to be reduced tozero.

[0125] For any extra cardiac support device, a volume compensatingchamber for the actuating volume is required. Since both the left andright sides are not completely independent of the other, the extracardiac support cannot alternatively pump the left and the right side.In general, for extra cardiac support, the two sides have to be pumpedsimultaneously. The artificial myocardium inherently has a volumereduction factor requiring only 25 cc of the hydraulic stroke volume fora biventricular support. This compensating chamber can be directlyincorporated on the energy converter as a flexing member or it can beseparated from the energy converter via a hydraulic conduit and shapedas a pancake flexible chamber. The latter is the preferred approachsince this provides additional flexibility in chamber placement. Thischamber would be a 2½″ diameter sac with flexible membranes to yield a 1cm excursion. The body of the chamber can be made from Angioflex®. Avelour layer can be the enclosure to provide tissue layer stability.

[0126] The control of the proposed artificial device needs toincorporate three key modes. These are, (1) synchronized contraction anddilatation with the heart when a normal R-wave is discerned, (2) nopumping during intermittent arrhythmia, and (3) pumping at ratesdetermined only by filling pressures during fibrillation and cardiacarrest. An additional design criterion is to ensure that the device doesnot work against the heart. The proposed control algorithm consists ofthree modes discussed below.

[0127] Synchronization is achieved by sensing the rhythm of the naturalheart, or paced signals for subjects with implantable pacers. Two basicapproaches are available, using either the P-wave or the R-wave, thechoice being governed by the conduction capability of the heart.

[0128] P-wave may be preferable as the reference for synchronization,since the right atrium remains free from any mechanical contact with thedevice. Naturally, if a subject suffers from frequent atrial flutter oratrial fibrillation, reliable P-waves would not be available. Inaddition, AV block would exclude the use of the P-wave. Patients withthese pathologic conditions would require R-wave sensing. Epicardialleads can be used for either sensing mode. For atrial sensing, a leadcan be sutured to the atrial appendage, while ventricular sensing can beachieved by a corkscrew electrode attached near the apex where no directsqueezing of the myocardium would occur. Bipolar electrode designs maybe used in order to localize signal reception especially, P-waves, withreduced noise pick-up in the acquired signals. Unipolar leads can beused for R-wave sensing since this is the simplest type of electrode forventricular epicardial fixation.

[0129] Whether P-wave or R-wave sensing is used, the algorithm isdesigned for synchronous contraction of the artificial device and themyocardium. For a subject with a regular heart beat, this can beachieved readily. For P-wave sensing, the device systole is timed toinitiate after ≈160 msec, a normal AV delay, following P-wave detection.With R-wave detection, device actuation is initiated immediately. Ananticipation algorithm which is based on the prior R-R intervals canalso be used. Such algorithms are available.

[0130] In FIG. 12A and 12B there is illustrated a subdermal port for theartificial myocardium assist system in case there is a failure of thehydraulic pumping capacity. FIG. 12A is a cross sectional view acrossthe skin interface 45 and FIG. 12B is an “x-ray” view of the systemviewed externally. In the case of a system failure, as a result ofelectronic or mechanical problems, the subdermal port can be accessedthrough a skin puncture with an array 48 of 15 gauge needles. Theprocedure would involve the extraction of the hydraulic fluid using a 50cc syringe. This extraction would collapse the artificial myocardiumcuff 11. A hand operated pneumatic pump (not shown) could then beconnected to the needle manifold 47 to activate the artificialmyocardium. The reason for the extraction of the hydraulic fluid andsubsequent manual use of a pneumatic pump is that the flow resistancethrough a 1 cm long parallel array of 15 gauge 1 mm ID) needles is lessthan 20 mm Hg for air, while the use of hydraulic fluid would result inpressure losses which are orders of magnitude higher. The artificialmyocardium system would be implanted through a median sternotomy. Thisprocedure is sufficiently simple so that it would be possible withoutbypasses, although severely compromised patients might require bypassfor support during the surgical procedure. Of course, other perhaps lessinvasive, surgical techniques could possibly be employed for thisimplantation. An appropriately sized artificial myocardium cuff 11 wouldbe wrapped around the natural heart. The energy converter 19 and thehydraulic reservoir 21 would be implanted in the thorax. It is estimatedthat the total volume and weight of the thoracic unit would beapproximately 105 cc and 165 g respectively. The energy converter andthe fluid reservoir, which in practice could be an integral part of theenergy converter, would be anchored to the rib cage with a flexiblehydraulic connection to the artificial myocardium cuff 11 and anelectrical cable tunneled through the costal diaphragmatic region to theelectronic components which could be implanted in the abdomen. Theseimplant locations are illustrated in FIG. 2. It is important that theartificial myocardium cuff 11 is anchored properly relative to thenatural heart such that during systolic contraction, the heart would notslip out of the myocardium cuff 11. Suitable attachment arrangementsare, for example, illustrated in U.S. Pat. No. 4,957,477.

[0131] The primary biocompatibility issue for the artificial myocardiumrelates to the epicardial tissue/cuff interface, or pericardialtissue/cuff interface. The material in contact with the epicardium orthe pericardium will be the polyetherurethane material Angioflex®. Otherimplanted components would consist primarily of cable jackets made fromeither medical grade room temperature vulcanizing rubber (RTV) orAngioflex® polyurethane. Non-flexing parts would consist of titanium ascasing for the energy converter and electronics packages. Infectiousrisks would be minimized in this design by the elimination ofpercutaneous exit and entry sites into the body, by quality control ofsurfaces and by choices of materials in contact with the tissue.

[0132]FIGS. 13 and 14 show the calculated pressure and volume ratios ofthe hydraulic and physiologic blood system fluids as a function of thenumber of tubes in the artificial myocardium. FIG. 13 shows thehydraulic physiological pressure ratio as a function of the number oftubes for both biventricular support and univentricular support. FIG. 14shows the physiological to hydraulic volume ratio. These figures showthat as the number of tubular segments increases, the required hydraulicstroke volume decreases while the required hydraulic pressure increases.For a volume ratio of three, the univentricular support (½ of the totalwrap) would require 11 segments, while the biventricular support wouldneed 23 segments because of the larger perimeter for biventricularsupport. This volume amplification between the hydraulic stroke volumeand the blood stroke volume is very significant, since it permits theactuating system for the artificial myocardium to be small and compactin size. In order to take advantage of this volume amplification, thepressure required to inflate the tubes to the appropriate extent for asignificant stroke work is approximately 10 times the afterload pressureof the blood being pumped. In the figures, the intersections of theuniventricular and biventricular supports with the dashed horizontalline indicate these operating points. FIG. 15 illustrates the bloodstroke volume in cc's as a function of the number of tube segments (n)in the artificial myocardium. As illustrated, the stroke volume does nothave a strong dependence on n, especially when n becomes large and thehydraulic displacement component becomes negligible compared with thecontractile effect. For a completely failed ventricle, in order togenerate 100 mm Hg of systolic pressure, the hydraulic drive pressurerequired will be approximately 1,000 mm Hg. This value is illustrated bythe star shown in FIG. 7. The benefits derived from operating the energyconverter 19 at lower flow and higher pressures are higher systemefficiency as a result of lower flow losses and smaller system size dueto the lower volume requirement. The gain in the hydraulic efficiencywill be primarily in the energy converter 19. For artificial myocardiumcuff 11, the flow velocities in the individual tubes is independent ofthe number of tubes. The volume of each tube scales with the area of thetube so that the flow velocity is a parameter determined only byphysiologic requirements. For a 65 cc stroke volume at a beat rate of140, the peak hydraulic flow velocity in the tubes is approximately 15cm per second, resulting in a dynamic pressure of approximately 0.1 mmHg, which has no impact on efficiency when compared to driving pressureson the order of one atmosphere. With this design, the wall stresses inthe artificial myocardium cuff 11 are, as in flow losses, independent ofthe number of tubes used in the wrap. The wall stresses in the wallsconnecting the tubes are only functions of the physiological parameters,such as the heart diameter and the physiological pressure. The wallstress is given as the product of the hydraulic pressure PH and the tuberadius, r, and is a constant, independent of the number of tubes, n.

[0133] The design consideration for the number of tubes per wrap will bedetermined by practical considerations such as the fabricationtechniques and energy converter efficiency.

[0134]FIG. 16 shows a mock loop used for in vitro tests of theartificial myocardium. Fluid from the reservoir 72 enters an atrium 68which empties through a inflow valve 70 to the ventricle 64. Theventricle consists of a cylindrical bladder which is surrounded byanother concentric cylindrical pouch. A space between the bladder andthe pouch is filled with a viscous fluid to simulate the ventricularwall. The exterior of the pouch has fitted eyelets spaced to accept anartificial myocardium simulating a left ventricular wrap. The outflowfrom the artificial ventricle 64 is coupled through another tri-leafletvalve 82 to an aortic compliance chamber 80, followed by a flow probe 78and a flow rotameter 76. The outflow resistance 74 is adjustable. Thereturn flow empties into the reservoir 72. For this study the artificialmyocardium assist system was actuated using a pneumatic drive consoleconsisting of a high pressure plenum and a low pressure plenum whichwere alternately switched to the device by solenoid valves initiatingsystole and diastole respectively. This drive mechanism replaces thehydraulic energy converter which would be employed in the implantablesystem. For this study FIG. 17 illustrates a linear relationship betweenthe afterload pressure (AOP) and driving pressures for the artificialmyocardium. The theoretical calculated value is shown as the solidcurve, while the square dots indicate the values determined in thisexperimental study. The flow output was maintained constant by adjustingthe aortic resistance. In the illustrated set of measurements the flowwas maintained at 6.5 liters per minute, with a filling pressure of 14.6mm of mercury, a beat rate of 169 beats per minute and a systolicduration and duty factor of 168 milliseconds and 40% respectively. Thedevice used in this study had 7 adjacent tubes and provided 50% wrappingof the pouch simulating the natural heart ventricle. The diameter andlength of the pouch were 6 cm and 5 cm respectively, fairly typical of asmall, left ventricle. Based on calculations, the anticipated strokevolume was 46.7 cc and the measured stroke volume was 38 cc, 82% of thetheoretical value.

[0135] A second set of measurements was obtained in this study bymaintaining a constant afterload, while the drive pressure was variedand the resultant ventricular flow recorded. The results of this studyare illustrated in FIG. 18. The solid curve shows the calculated valuesand the set of square points illustrate the measured values. Theconditions were similar to those for the experiment illustrated in FIG.17. FIG. 18 also illustrates the calculated flow versus the drivepressure relationship. The outflow pressure was set at 115 mm Hg, whichis the intercept of the drive pressure at zero flow. This study showedthat the experimental pneumatic drive pressure was slightly higher thanthat which would have been predicted by the theoretical calculation. Thedata from this study indicates by controlling the drive pressure, whichis equivalent to adjusting the contractility of the artificialmyocardium, both flow and pressures can be enhanced.

[0136] In FIGS. 19A and 19B there is illustrated one embodiment of agirdle for wrapping around a heart to constrain dilatation of theventricle and limit the amount of energy and oxygen required to maintainthe heart muscle in tension.

[0137] In FIG. 19A the natural heart 110 is shown with the leftventricle 112 somewhat dilated and with a girdle 117 surrounding boththe left ventricle 112 and the right ventricle 111. The girdle 117 isformed as illustrated in FIG. 20, with a series of horizontal segments113 a-113 i encircling the heart 110, the segments toward the apex ofthe heart being smaller in cross section and in length. The girdlesegments 113 are filled with hydraulic fluid which is maintained at aconstant volume during the beating of the heart. In this arrangement thegirdle is entirely passive and a distensible girdle lining 118 conformsto the shape of the heart at the myocardium girdle lining interface byvirtue of the pressure of the fluid filled segments 113 against thedistensible inner lining 118. As shown in FIG. 3, when this girdle isimplanted around a natural heart the volume is controlled through athree-way valve 127 which controls the amount of fluid supplied to thegirdle segments 113 from reservoir 125, which is formed of a rigidcasing 124.

[0138] According to equation (1), it can be seen that an increase inmechanical work by a large factor results in a small increase in oxygenconsumption, but an increase in tension time causes a large increase inoxygen consumption. Passive girdling of the heart, as illustrated inFIGS. 19-21, acts to limit or reduce the ventricular size of thediseased ventricle. Over an extended period of time, which may be daysor weeks, the fluid 114 volume may be increased, thereby decreasing theperiphery of the interface lining 118 of the girdle, which may over aperiod of time actually decrease the dilatation of the ventricle 112.

[0139] In FIG. 21 a control system for controlling the fluid pressure inthe segments 113 according to the tension in liner 118 is shown. Thefluid pressure in girdle 113 is controlled by a feedback loop includinga strain gauge 142 placed at the interface between the inner lining 118and the myocardium providing a sensed value for the tension of themyocardium, to hydraulic actuating electronics 122 which may be aconventional hydraulic control circuit. The electronic actuator 122controls a conventional mechanical fluid actuator 123 which provides forincrease or decrease of fluid within the girdle 117. This actuatoroperates in conjunction with a three-way valve 127 and fluid reservoir125. The change in volume effected by this feedback, is not intended to,nor does it operate in the time frame of the beating of the naturalheart. It is meant to adjust the volume over a much longer time period,typically days, weeks or months.

[0140] In this configuration, the series of generally cylindricalsegments 113 are typically formed of non-distensible material. They areattached to one another along the long axis of the cylinder and may befilled with fluid either individually or in parallel. When the fluidvolume within the compartments 113 is very low, then the girdle 113assumes the shape shown in FIG. 19A providing for a large innerdiameter. On the other hand, when the fluid volume is increased thesegments assume, at fill inflation, a circular cross section therebydecreasing the inner perimeter very substantially, as illustrated inFIG. 19B. Thus, by controlling the volume of the fluids supplied to theindividual segments 113, the inner diameter of the girdle 117 can beadjusted to be a close fit to the natural heart. This configuration hasthe advantage that, since there is no single vertical compartment, thereis no gravity pooling of fluid in one portion of the girdle 117. FIG. 22illustrates a second embodiment of this invention. The girdle 130 ofFIG. 22 is an adjustable girdle made from a synthetic material that canlimit tension, but is otherwise deformable to conform to the anatomicalgeometry of the heart. In this case, the girdle 130 is formed of aconfining net 132 which is wrapped around the heart from the apex to theatrioventricular (A-V) groove. The purpose of this net is to limit themaximum diastolic dimension of the heart, while offering no resistanceto systolic ejection. In the design illustrated in FIG. 22 a number ofinterlinked two-dimensional loops such as lightweight plastic rings 133are interconnected to form the girdle or wrap 130. The loops 133 arefree to move in all directions without restraint, since none arephysically connected to each other. Rather, they are interlocked byhaving the loops or rings 133 pass through one another. The design ofFIG. 22 presents no systolic load to the contracting heart. Theloop-mesh 132 can readily conform to the shape of the heart with thechange in surface area accompanying the heart contraction readilyaccommodated by the free loops.

[0141] An alternative form of this loop-mesh girdle is shown in FIG. 23.In FIG. 23 a string system 134 is included with the string attached tothe loops 133 to effect change in the size of the mesh by virtue ofpulling the strings. This arrangement is able to accommodate a treatmentmodality for scheduled size reduction to the heart over a suitableperiod of time. In FIG. 23, a segment of the girdle or wrap 130 isshown. The original size of the wrap can be seen at the wide edge 136,while the narrowed down section is seen at the ridge 138 of the wrap.Pulling on the two ends of two sets of strings reduces the size of themesh in two directions. This can be done during a thoracoscopy orthrough a cutaneous access port. In the construction illustrated inFIGS. 22 and 23 the net 130 will be attached at several attachmentpoints, typically 4 to 6 in number, at the A-V groove and also perhapsnear the apex of the natural heart. At the original implant the surgeonwill optimize the fit to the heart as it is existing and will adjust thesize through the mechanism described above. This design will accommodatespontaneous heart size reduction even though some parts of the mesh mayadhere to the epicardium. However, due to relative motion between theloops, it is unlikely that the mesh will become fully encapsulated. InFIG. 24 there is shown a girdle in accordance with this invention whichis formed of a sheet of an expanded polytetrafluroethylene (PTFE)material 124, prestressed such that it remains below its elastic limitand its tension in the plane of the sheet is sufficient to createradially inward forces, thus resisting expansion while permitting inwardcompression. In other words the girdle will resist further expansionwhile fittingly accommodating shrinkage. Other materials may beemployed, provided that they exhibit the above elasticitycharacteristics.

[0142] In FIG. 25 there is illustrated a cross sectional view of atissue engineered girdle lining having a polymer scaffold 131 which hasbeen seeded with myocardial cells harvested from the recipient mountedon a polymer substrate 131, the substrate either facing a girdlestructure or forming the inner surface of that girdle. The tissueengineered lining faces the patient's myocardium. Such a lining reducesthe irritation which may occur between the epicardium and artificialmaterials employed to form the girdle itself. The lining 130 would, overtime, integrate biologically to the patient's myocardium.

[0143] Techniques for cell scaffold engineering are described in theliterature. Two examples being, Biodegradable Polymer Scaffolds forTissue Engineering by Lisa E. Freed, Gordana Vunjak-Novakovic, Robert J.Biron, Dana B. Eagles, Daniel C. Lesnoy, Sandra K. Barlow and RobertLanger and Tissue by Robert Langer and Joseph P. Vacanti, Biotechnology,Vol. 12, July 1994 and Tissue Engineering, Robert Langer and Joseph P.Vacanti, Science, Vol. 260 May 14, 1993.

[0144] This tissue engineering techniques may also be employed withrespect to other artificial materials which come in contact with theheart in various surgical situations including the active devicesdescribed hereinabove and in U.S. patent application Ser. No.08/490,080, filed Jun. 13, 1995.

[0145] Having described the above specific embodiments of thisinvention, other embodiments implementing the concepts of this inventionwill doubtless occur. While specific details of an artificial myocardiumand artificial myocardium assist system have been illustrated, it willbe understood that other embodiments may be formed employing theprinciples of this invention.

1. A method for treatment of a patient, whose heart is characterized byventricular dilatation comprising the steps of, wrapping a girdle aroundat least the ventricle of said patient's heart; and maintaining saidgirdle in a passive state for an extended period of time, said girdlebeing formed such that it can adjust in size and shape to conform to theouter shape of said ventricle and to not expand its dimension in adirection away from said natural heart.
 2. A method in accordance withclaim 1 wherein said girdle is formed of a sheet of material prestressedin the plane of said sheet to a value below the elastic limit of saidmaterial, said sheet having a tension which limits extension away fromsaid heart, while providing compression forces radially inward towardsaid heart.
 3. A method in accordance with claim 1 wherein said girdleis formed of a vertically oriented series of sealed, independent,generally horizontally extended cylindrical segments, said verticalorientation being parallel with an axis of the heart running to itsapex, including the further step of, introducing fluid into saidcylindrical segments to decrease the inner perimeter of said girdle sothat its size conforms generally to the size of said patient'sventricle.
 4. A method in accordance with claim 3 and further includingmeans for increasing the volume of fluid within said sealed volume in acontrolled fashion to decrease the dimensions of the inner lining whensaid ventricle decreases in size over an extended period.
 5. A method inaccordance with claim 1 wherein said girdle is formed of a net ofinterlocking loops, unattached to one another.
 6. A method in accordancewith claim 5, wherein and said interlocked loops are furtherinterconnected by strings extending in at least a first dimension andincluding the further step of pulling on said strings to decrease thedimension of said girdle in position around said patient's heart.
 7. Amethod in accordance with claim 3 and further including the step of,placing an inner lining between said cylindrical segments and the outersurface of said patient's heart and, placing a tension measuring sensorat the interface between the outer surface of said patient's heart andthe inner lining of said girdle, said sensor providing an output signalindicative of the tension of said lining adjacent to said sensor, and,providing said output signal to means for adjusting the amount of fluidwithin said cylindrical segments until the tension at said lining is ata predetermined value.
 8. Apparatus for providing passive support to anatural heart characterized by ventricular dilatation comprising, agirdle for wrapping around said heart, said girdle being formed of asheet of material prestressed in the plane of said sheet to a valuebelow the elastic limit of said material, such that said sheet has atension which limits extension away from said heart, while providingcompression forces radially inward toward said heart.
 9. Apparatus forproviding passive support to a natural heart characterized byventricular dilatation comprising, a girdle for wrapping around saidheart, said girdle comprising a net of interlocked plastic loops, notattached to one another.
 10. Apparatus for providing support to a heartcharacterized by dilatation of a ventricle comprising, a girdle formedof a vertically oriented series of sealed, independent, generallyhorizontally oriented cylindrical segments, said vertical orientationbeing parallel with an axis of the heart running to its apex, means forproviding a fluid to said segments to control the dimensions of saidgirdle, and means for changing the volume of fluid provided to saidsealed volume only after an extended period of time.
 11. Apparatus inaccordance with claim 10 and further including a distensible liningformed between said cylindrical segments and said natural heart. 12.Apparatus in accordance with claim 10 and further including means forproviding as an interface between said inner lining and the outersurface of said heart, a cellular wall constructed of myocardial cellsharvested from said heart and mounted on a scaffold to generate saidcellular wall.
 13. Apparatus for providing support to a heartcharacterized by dilatation of a ventricle, comprising a girdle formedwith open ends, the wall of said girdle being formed of aninterconnected vertically oriented series of closed, horizontallyoriented, tubes constructed of a flexible, nondistensable material,forming a fluidically sealed volume, said vertical orientation beingaligned with an axis of the heart running to its apex, means forcontrollably filling said tubes with a fluid, such that when completelyfilled, each of said tubes assumes a tubular cylindrical shape with acircular cross section, and when emptied of said fluid, each of saidtubes assumes a collapsed shape, and means for providing a selectedamount of filling to said tubes with said fluid, and changing saidamount only in response to decrease of said dilatation after an extendedperiod of time.
 14. A method of generating the interface between theinterior of an external girdle for a natural heart and the myocardium ofsaid natural heart comprising the step of, providing a scaffold ofbiologically inert material between the lining of said girdle and theexterior myocardial surface of said heart and generating a wall on saidscaffold by application of myocardial cells to said scaffold, saidmyocardial cells being harvested from said heart.